System and method for creating a stable optical interface

ABSTRACT

A system and a method for creating a stable and reproducible interface of an optical sensor system for measuring blood glucose levels in biological tissue include a dual wedge prism sensor attached to a disposable optic that comprises a focusing lens and an optical window. The disposable optic adheres to the skin to allow a patient to take multiple readings or scans at the same location. The disposable optic includes a Petzval surface placed flush against the skin to maintain the focal point of the optical beam on the surface of the skin. Additionally, the integrity of the sensor signal is maximized by varying the rotation rates of the dual wedge prisms over time in relation to the depth scan rate of the sensor. Optimally, a medium may be injected between the disposable and the skin to match the respective refractive indices and optimize the signal collection of the sensor.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates generally to stabilizing an opticalinterface and, more specifically, to creating a reproducible and stableoptical interface between biological tissue and an optical blood glucosesensor.

2. Related Art

Monitoring of blood glucose concentration levels has long been criticalto the treatment of diabetes in humans. Current blood glucose monitorsinvolve a chemical reaction between blood serum and a test strip,requiring an invasive extraction of blood via a lancet or pinprick.Small handheld monitors have been developed to enable a patient toperform this procedure anywhere, at any time. But the inconvenience ofthis procedure—specifically the blood extraction and the use anddisposition of test strips—has led to a low level of compliance. Suchlow compliance can lead to serious medical complications. Thus, anon-invasive method for monitoring blood glucose is needed.

Studies have shown that optical methods can detect small changes inbiological tissue scattering related to changes in levels of bloodsugar. Although highly complex, a first order approximation ofmonochromatic light scattered by biological tissue can be described bythe following simplified Equation 1:I _(R) =I _(O) exp [−(μ_(a)+μ_(s))L]  Eq. 1where I_(R) is the intensity of light reflected from the skin, I_(O) isthe intensity of the light illuminating the skin, μ_(a) is theabsorption coefficient of the skin at the specific wavelength of light,μ_(s) is the scatter coefficient of the skin at the specific wavelengthof light, and L is the total path traversed by the light. From thisrelationship, it can be seen that the intensity of the light decaysexponentially as either the absorption or the scattering of the tissueincreases.

It is well established that there is a difference in the index ofrefraction between blood serum/interstitial fluid (blood/IF) andmembranes of cells such as blood cells and skin cells. (See, R. C.Weast, ed., CRC Handbook of Chemistry and Physics, 70th ed., (CRCCleveland, Ohio 1989)). This difference can produce characteristicscattering of transmitted light. Glucose, in its varying forms, is amajor constituent of blood/IF. The variation of glucose levels inblood/IF changes its refractive index and thus, the characteristicscattering from blood-profused tissue. In the near infrared wavelengthrange (NIR), blood glucose changes the scattering coefficient more thanit changes the absorption coefficient. Thus, the optical scattering ofthe blood/IF and cell mixture varies as the blood glucose level changes.Accordingly, an optical method presents a potential option fornon-invasive measurement of blood glucose concentration.

Non-invasive optical techniques being explored for blood glucoseapplication include polarimetry, Raman spectroscopy, near-infraredabsorption, scattering spectroscopy, photoacoustics and optoacoustics.Despite significant efforts, these techniques have shortcomings such aslow sensitivity, low accuracy (less than current invasive home monitors)and insufficient specificity of glucose concentration measurement withinthe relevant physiological range (4-30 mM or 72-540 mg/dL). Accordingly,there is a need for an improved method to non-invasively monitorglucose.

Optical coherence tomography, or OCT, is an optical imaging techniqueusing light waves that produces high resolution imagery of biologicaltissue. OCT creates its images by focusing a beam of light into a mediumand interferometrically scanning the depth of a linear succession ofspots and measuring the absorption and/or the scattering of the light atdifferent depths in each successive spot. The data is then processed topresent an image of the linear cross section of the medium scanned. Ithas been proposed that OCT might be useful in measuring blood glucose.

One drawback associated with using OCT for monitoring blood glucose isthe signal noise associated with optical interferometry, also known asspeckle. As discussed in U.S. application Ser. No. 10/916,236 by M.Schurman, et al, entitled “Method and Apparatus for Monitoring GlucoseLevels In A Biological Tissue,” to reduce speckle, a glucose monitorincorporating OCT methodology may scan a beam of collimated lightcontinuously and laterally across a two-dimensional surface area of apatient's tissue or skin, while interferometrically scanning the tissuein depth. Preferably, the scanning is accomplished with a small,lightweight, and robust mechanism that can be incorporated into a sensorto be used in a fiber-optics based product or, alternately, a nonfiber-optics based product. One main objective of using this type ofsensor is to generate a reproducible stable optical interface betweenthe subject's skin and optical path of the sensor in order to takemultiple readings from the same lateral location on the skin whilemaintaining the integrity of the optical interface. As discussed below,there are multiple problems associated with providing and maintaining astable and reproducible optical interface between an OCT sensor and theskin of a patient.

Two Basic Optic Designs

Two well known sensor designs that use OCT are schematically shown inFIGS. 1 and 2. FIG. 1 shows a design based on the use of two rotatingwedge prisms to change the angle of collimated light incident on afocusing lens. In FIG. 1, incoming light beam 101 hits a collimatinglens 102, which splits the beam 101 into multiple parallel beams oflight, or collimated light 103. The collimated light 103 then passesthrough one or more wedge prisms 104, which are rotating at predefinedrates. As shown in FIG. 1, dual rotating wedge prisms 104 generate anangular deviation in the collimated light 103 from the optical axis ofthe sensor, which is the “centerline” axis passing through the elementsof the sensor, perpendicular to the surface area of skin 109 to betested. By deviating the angle of the collimated light 103, the focalpoint of the light moves around on a focal plane of an optical window108 that is flush against the skin 109, thereby scanning differentlateral locations on the skin 109. As shown at 105, once passing throughwedge prisms 104, the parallel rays of collimated light 103 may beangled away from the optical axis, depending on what portion of thewedge prisms 104 the collimated light 103 passes through. The angledbeams 105 then pass through a focusing lens 106, and begin to focustogether to a focal point 107 at the bottom surface of an optical window108.

FIG. 2 shows a similar concept to FIG. 1, however the dual wedge prisms104 of FIG. 1 are replaced with an angled mirror 201, for example, a 45degree angled mirror, that oscillates along two axes, thereby deviatingthe angle of collimated light 103 from the optical axis in order to movethe focal point 108 around on the surface area of skin 109. Accordingly,this OCT sensor design is well known in the art. Both designs facilitatescanning an area of skin by deviating the angle of collimated beam 103from the optical axis, thereby moving the focal point 107 a proportionaldistance laterally in the focal plane along the bottom of the opticallens 108, and, accordingly, along the surface area of the patient's skin109.

While both sensor designs provide mechanisms for incorporating OCT intoa noninvasive blood glucose sensor, there are several drawbacksassociated with the above designs as described below.

Variations in Optical Path Length

One drawback associated with the dual wedge prism sensor design of FIG.1 is illustrated in FIG. 3. In an interferometer, the optical pathlength of a beam of light is determined by the physical or geometricpath length of the beam and the index of refraction of the medium whichthe beam is passing through as shown in Equation 2:L _(OPT) =n·L _(GEO)  Eq. 2where “L_(OPT)” is the optical path length, “n” is the index ofrefraction, and “L_(GEO)” is the geometric or physical path length

As shown in FIG. 3, depending on the position of the wedge prisms 104 atthe time the collimated beam 103 shines through, while the geometricpath length of the collimated beam 103 stays the same, the index ofrefraction changes due to the changing thickness of the wedge prisms 104as the prisms rotate, thereby altering the optical path length of thecollimated beam 103. This continuous change in the thickness of thewedge prisms 104 continuously alters the optical path length of thecollimated beam 103 as it passes through. As shown in FIG. 3, theplacement of the wedges may extend the length of the optical path,making it seem as though the skin 109 is moving away from the sensor.Thus, three optical scans taken through the dual wedge prisms 104 whenthe prisms 104 are in different rotated positions produce three scansbeginning at different positions in depth. Since the sensor data is anaverage of multiple scans, if each scan begins at a different positionin depth, the resulting ensemble average will not be representative of atrue averaging of multiple scans.

For example, in FIG. 3, when the collimated beam 103 passes through thethinnest area of the wedge prisms 104, as shown at 301, the sensorbegins to collect data at Depth A, interpreting the interface betweenthe optical window 108 and the skin 109 to be at Depth A, as shown at302. However, when the collimated beam 103 passes though a thin portionof the first wedge prism and a thick portion of the second wedge prism,as shown at 302, the sensor begins to collect data at Depth B,interpreting the interface between the optical window 108 and the skin109 to be at Depth B, as shown at 304. Further, when the collimated beam103 passes through the thickest portion of both wedge prisms, as shownat 305, the sensor begins to collect data at Depth C, interpreting theinterface between the optical window 108 and the skin 109 to be at DepthC, as shown at 306. Since typically multiple scans (e.g., greater than100 scans) are taken and then averaged to reduce speckle, scans taken atdifferent positions in depth cannot be averaged. Thus, a solution tothis problem is desired.

Another drawback associated with the dual wedge prism sensor is thedistortion of the scan along the depth axis or z-axis of the light beamentering and exiting the skin. If the rotation speed of the wedge prisms104 is several orders of magnitude larger than the depth scan rate ofthe optical sensor, then the depth scale measured by the scan is either“stretched” or “shrunk” by the entire amount of the difference inoptical path induced by the changing thickness of the wedge prisms 104.However, if the rotation speed of the wedge prisms 104 is much slowerthan the depth scan rate, then the changing thickness of the wedgeprisms 104 has a minimal effect on the depth scale. For example, if thedepth scans occur at 60 Hz, which means that the sensor completes onedepth scan within in 1/60^(th) of a second, and the prisms rotate at3600 rpm, then each wedge prism makes a full rotation during the time ittakes the sensor to complete one depth scan. Because the thickness ofeach wedge prisms varies as the prisms rotate, the optical path lengthchanges during each depth scan, which distorts the depth data collectedby the sensor by changing the depth scale during a single scan. Thus,there is an optimization that must occur between the depth scan rate andthe prism rotation rate such that the entire surface area is thoroughlyscanned while minimizing the z-axis scan distortion.

Scan Pattern Stability

Accordingly, it is desired is that each depth scan be taken at adifferent lateral position on the surface of the skin 109 such that theensemble of all the depth scan positions are randomly and uniformlydistributed throughout the scan region. The lateral locations of eachdepth scan must be spatially independent to 1) effectively encompassregions of blood glucose change during a sensor reading and 2)effectively reduce speckle. However, a problem associated with the dualwedge prism sensor in FIG. 1 and the oscillating mirror sensor in FIG. 2is the inability to capture each depth scan position due to the angularvelocity of the wedge prism(s) 104 or the oscillation rate of the angledmirror 201 being harmonic in phase with the depth scan rate of theoptical sensor, i.e., the frequency of the angular velocity is amultiple or integral of the depth scan rate of the sensor. When eitherthe angular velocity or oscillation rate is an integral of the depthscan rate, the two rates “beat” against each other, and produce a lossof conformal coverage of the surface area of the skin 109 being scanned.

As shown in FIG. 4B, when using a single rotating wedge prism oroscillating angled mirror in a sensor as described above, the optimalresult of multiple depth scans is a circle pattern on the surface areaof the skin 109, which each “dot” representing a depth scan. Each depthscan occurs along the path of this circle pattern, effectively breakingthe circle up into a series of scanned points. However, if the angularvelocity is an integral or harmonic of the depth scan rate, the depthsscans begin to overlap in location, thereby producing an incompletecircle pattern and a loss of spatially independent depth scans, as shownin FIG. 4A. With an overlap of depth scans, the same locations of tissueare scanned, causing less speckle reduction and poor imaging ofstructures within the scanned tissue. The problem becomes even morepronounced in the case of a sensor with two wedge prisms, as shown inFIG. 4C.

Focal Plane Instability

Another challenge presented by both the wedge prism design in FIG. 1 andthe oscillating mirror design in FIG. 2 is the inability to maintain thefocal point 107 of the focused collimated beam on the focal plane, orthe interface between the optical window 108 and the surface area of theskin 109 being scanned. Optical lenses do not project an image onto aflat plane, such as the flat bottom surface of the optical window 108,but, instead, naturally project an image onto a curved surface, muchlike the curved interior of the eye. This curved surface is well knownas a Petzval surface. Thus, as the collimated light 103 enters thefocusing lens 106, the focal point 107 of the collimated light 103traces out a curved focal plane or Petzval surface based on the designof the focusing lens 106, caused by the angular deviation from theoptical axis due to the wedge prisms 104 in FIG. 1 or the angled mirror201 in FIG. 2. Thus, the flat bottom of the optical window 108 does notallow the focal point 107 to remain on the focal plane.

When the focal point 107 moves off of the Petzval surface, theefficiency of the focused light being collected begins to drop, sincefocal plane is where the light capture is maximized. Additionally, thedepth scale of the focused light is affected such that the displacementof the focal point 107 off of the focal plane results in an equivalentloss in the depth scale of the signal. This results in a blurring of theoptical axis, causing measurable details within the skin to be blurredor washed out. Thus, a displacement of the focal point off the focalplane results in a reduction in the sensor signal intensity and ablurring of the optical axis.

Additionally, optical lenses are not perfect. Therefore, as the focalpoint 107 moves away from the optical axis due to the rotating wedgeprisms 104 or the oscillating angled mirror 201, the focused beam driftsaway from the skin 109 and back towards the focusing lens 106, and,thus, moves off the focal plane. As discussed above, when the focalpoint 107 is no longer on the focal plane, the collection efficiency ofthe light drops, resulting in the collected data incorrectly indicatinga reduction in power. This, in turn, alters the depth of the focusedbeam, thereby unwittingly washing out details in the skin and loweringthe resolution and integrity of the scan.

Skin/Sensor Optical Interface

The surface of the skin is “rough” relative to the light entering andexiting the skin during an optical scan. This is well known as opticalroughness. Additionally, the refractive index of the skin being scannedtypically is different from the refractive index of the material of anoptical window of a sensor. As shown in FIG. 5A, the optical window 503is not necessarily flush against the surface of the skin 504, due tooptical roughness 505 of the skin. Accordingly, as incident light 501 isdirected towards the skin, some of the light is reflected and/ordiffracted, as shown at 502, because there is a mismatch between theindex of refraction of the optical window 503 and the index ofrefraction of the skin 504. This mismatch of refractive indices and, inaddition, the space between the skin 504 and the optical window 503 dueto the optical roughness 505 reduces the reliability of data taken bythe sensor.

FIG. 5B displays two scans taken at the same location on the skin butmeasured at different points in time with constant optical contactbetween the skin 109 and the optical window 108 of a sensor. Such scansmay be produced by either the dual wedge prism sensor of FIG. 1 or theangled mirror sensor of FIG. 2. Data line 506 represents an averagedoptical scan taken at Time 0 while data line 507 represents an averagedoptical scan taken thirty minutes after Time 0. Typically, the focusedbeam hits the interface between the optical window 503 and the skin 504,a sharp rise or peak in the signal is produced, as shown at peaks 510and 511. The signal then drops as the beam moves through the skin 504and begins to rise again as the beam hits the interface between theepidermis and dermis layers, as shown at peaks 508 and 509. The signalagain drops and continues to drop as the beam reaches the desired depththen returns back to the sensor.

As shown in FIG. 5B, while constant optical contact is maintainedbetween the skin 504 and the optical window 503 of the sensor, over timethe optical signal drifts, as illustrated by the peaks at the interfacebetween the dermis and epidermis layers, which rises over time, frompeak 508 at Time 0 to peak 509 at Time 0+30 minutes. However, the peakat the interface between the optical window 503 and the skin 504 dropsover time, from peak 510 at Time 0 to peak 511 at Time 0+30 minutes.This change in signal intensity is due to a gradual change in theoptical interface created by an accumulation of sweat and skin oils atthe interface of the optical window 503 and the skin 504, as shown at512 in FIG. 5C, which serves as an optical transition for the incidentlight 501 to efficiently travel from the optical window 503 to the skin504. Additionally, the accumulation of sweat and skin oils smoothes outthe optical roughness of the skin. Although the refractive index betweenthe optical window 503 and the skin 504 will stabilize or reach anequilibrium value due to sweat, oil, and other fluids produced by theskin over time, this process could take upwards of 60-90 minutes.Unfortunately, these changes in signal intensity over this extendedperiod of time may completely mask the changes that are occurring alongthe OCT signal, and thus prevent proper correlation of changes in theOCT signal to changing glucose levels, as discussed in U.S. ProvisionalApplications Nos. 60/671,007 and 60/671,285, both entitled “Method ForData Reduction and Calibration of an OCT-Based Blood Glucose Monitor.”Thus, multiple scans taken over time cannot produce a reliablemeasurement from the same lateral location on the skin. In addition, apatient would be required to place the sensor onto his or her skin andwait 60-90 minutes before using it, in order to receive reliable andreproducible results, which creates an inefficient sensor.

Thus, a need exists for an optical sensor for measuring blood glucoselevels and other physiological effects that overcomes the deficienciesdiscussed above.

SUMMARY OF INVENTION

According to one embodiment of the present invention, a system forgenerating a stable and reproducible optical interface includes anOCT-based interferometer connected to an optical sensor that utilizes acollimated beam of light and comprises dual wedge prisms to move thecollimated beam to different lateral locations on the skin, and adisposable optical lens apparatus that attaches to the skin surfaceusing an adhesive, where the disposable optical lens apparatus comprisesa focusing lens and an optical window that interfaces directly with theskin. Alternately, the optical sensor may utilize an angled mirror thatoscillates along two axes to move the beam of light to different laterallocations on the skin surface.

By using a disposable optical lens apparatus, a patient may place thesensor onto the optical lens apparatus, take a reading, then remove thesensor and leave the optical lens apparatus attached to his or her skin,for example, on an arm. When another reading is taken at a later time,the patient simply reattaches the sensor to the optical lens apparatus,guaranteeing that the lateral location of the sensor remains the same,in order to produce a comparable optical scan. At some point in time,the patient may remove the disposable optical lens apparatus and discardit, only to replace it with another. Thus, the disposable optical lensapparatus may be made from different materials, such as, for example,glass, plastic, or other polymer material, and may be customized foreach patient's needs. A computer also may be connected to the opticalsensor and/or interferometer, where the computer manipulates the sensordata and produces physiological data, such as blood glucose levels.

As mentioned above, multiple scans may be taken during a single sensoruse and then averaged together to reduce or remove the speckleassociated with an OCT-based system. To account for variations in theoptical path length of the collimated beam produced by the varyingthicknesses of the rotating dual wedge prisms, the resulting scan datais manipulated. According to an embodiment of the present invention, amethod for resolving the variations in optical path length includes thesteps of (i) locating the first peak, which represents the interfacebetween the optical window and the patient's skin, of the first scantaken by the sensor, (ii) locating the first peak in each subsequentscan taken during the single use, and (iii) normalize each first peak inthe subsequent scans against the peak of the first scan. The methodfurther comprises the step of (iv) averaging the normalized scans toproduce an averaged scan result. To locate the peaks, algorithms such asGaussian peak fitting and second-derivative residual methods may be usedand are well known within the field of the invention.

An alternate embodiment of the present invention presents a moretime-efficient method for resolving the variations in the optical pathlength. The method includes the steps of (i) setting a peak thresholdtrigger in the signal intensity and (ii) holding off of true dataacquisition until the signal hits the threshold trigger. Once signalreaches the threshold trigger, the system begins to collect the scandata. Different optical arrangements may require different thresholdtriggers, where optical arrangements may vary due to the angle of thewedge prisms in the optical sensor. However, to optimize the thresholdtrigger, at least a 10 db difference may exist between the thresholdtrigger and the first peak intensity value, where the signal intensityis measured in decibels. For example, if the first peak measures 60 db,then the threshold trigger is set to less than or equal to 50 db.Additionally, the threshold trigger may be set above the highest noisepeak produced by the signal until the focused beam hits the opticalwindow, where the signal begins to rise in intensity. For example, ifthe highest noise peak is 30 db and the first intensity peak reaches 60db, then setting a threshold trigger between 30 db and 50 db ispreferable. Since the most useful data is acquired beginning typicallyaround 150 microns in depth (within the dermis layer of the skin), andthe first peak in intensity typically occurs around 20 or 30 microns indepth, by setting a threshold trigger near the rise of the first signalpeak, any mismatch in the optical path length will be less than half thecoherence length of the optical sensor system, which is below theresolution of the interferometer.

The coherence length of the optical sensor system, which is a measure ofthe depth resolution of the system, is broadly inversely related to thebandwidth of the optical source of the system, such as, for example, asuperluminescent diode. Thus, as the bandwidth of the optical sourceincreases, the coherence length of the system decreases, andaccordingly, the depth resolution of the system improves. The interfacebetween the optical sensor and the skin has a specific peak intensityvalue, for example, 60 dB, and the width of the peak is the coherencelength of the optical sensor system, for example, 30 microns. However,for each depth scan, the optical sensor/skin interface peak doesn'talways occur at the exact location in depth, i.e., the peak location maybe offset by a few microns in depth. If, for example, the thresholdtrigger is set to a value that is near the signal peak intensity value,then the offset of the location of each peak value for each depth scancannot be more than a fraction of the coherence length, which is belowthe resolution of the optical system. Thus, the offset does not affectthe data collected by the sensor and the depth scans may be averaged toreduce speckle and to produce an accurate sensor reading.

According to an aspect of the embodiment, the optical sensor system maybe set to acquire data once the focused beam reaches a specificstructural feature. For example, the threshold trigger may be set tocorrespond to an intensity value of light once the focused light reachesthe interface between the skin and the optical window, which may occur,for example, at a depth of one-half of a millimeter (“mm”). Thus, if theoptical window/skin interface occurs at an intensity value of 60 dB,then the trigger threshold may be set to a value of 50 dB. Therefore,the optical window/skin interface becomes a reference point for eachdepth scan to be lined up against, in order for the depth scans to beaveraged.

According to another embodiment of the present invention, a method forminimizing the distortion in the depth scale due to change in thicknessof the dual wedge prisms as they rotate includes the step of optimizingthe depth scan rate versus the prism angular velocity in order tominimize any distortion of the scan in depth, or along the z-axis. Ifthe depth scans occur at a rate at or near the angular velocity of thewedge prism, then each depth scan performed by the sensor occurs withina time period close to the time period of a single rotation of the wedgeprisms. As discussed above, because the wedge prisms are not a uniformthickness and the thickness affects the refractive index and the opticalpath length, as the prisms rotate, the depth of each depth scan isdistorted within a single scan because the optical path length ischanging during a single scan when the time periods are close or exact.To prevent this problem, the method includes the step of setting theangular velocity of the wedge prisms to a value such that the lateralposition of the scan spot on the skin surface moves a distance that isless than ten times (“10×”) a diameter of the scan spot during the dataacquisition of a single depth scan. This method allows the optical pathlength to remain stable during each depth scan taken.

In yet another embodiment of the present invention, a method forstabilizing the scan pattern of the optical sensor includes the step of(i) setting the angular velocity of the wedge prisms to a non-harmonicphase value in relation to the depth scan rate. By doing so, conformalcoverage of the scanning area may be achieved. However, due to the driftof the angular velocities common in such a system, it is likely that theangular velocity will drift into a harmonic phase of the depth scanrate, and conformal coverage will be lost. Thus, the method furthercomprises the steps of (ii) varying the angular velocities of the dualwedge prisms during the total time of an entire sensor reading (i.e.,1500 scans), and (iii) varying the angular velocities of each wedgeprism with respect to the other wedge prism over the total time of thesensor reading. By varying both the angular velocity of the wedge prismsover time in relation to the depth scan rate, and the angular velocityof each wedge prism over time in relation to the other wedge prism,conformal coverage of the scan surface area is maximized. According toan aspect of the present embodiment, the method may be modified to varythe oscillation rate of the angled mirror in the mirror sensor such thatthe oscillation rate in both axes of movement is not a harmonic of thedepth scan rate of the sensor.

According to an alternate embodiment of the present invention, in anoptical sensor with rotating dual wedge prisms, two harmonically relatedphase signals may be used to vary the angular velocities of each wedgeprism so long as the time period of one phase signal associated with oneof the wedge prisms is several times longer than the time period of onephase signal associated with the other wedge prism, and both phasesignals are non-harmonic values of the depth scan rate. For example, if2.0 Hz and 0.02 Hz are the angular velocities maintained over time ofthe wedge prisms, and the depth scan rate is 57 Hz, the problem isminimized and conformal coverage of the scan pattern is maximized. Theembodiment encompasses numerous ways to vary the angular velocity of thewedge prisms, for example, a saw tooth wave, a sinusoidal wave, atriangle wave, etc.

In yet another embodiment of the present invention, a method foroptimizing an amount of light entering and exiting an area of skinincludes modifying the disposable optical lens as described above byincorporating a dome shape to the bottom surface of the optical window.The dome shape is designed to represent the Petzval surface of thefocusing lens, and follows the variation in the focal point displacementthat occurs as the focal point deviates from the optical axis throughincreasing incidence angles of the focused beam. Thus, the Petzvalsurface rests between the skin and the optical window of the disposable.Additionally, the Petzval surface also improves the interface betweenthe disposable apparatus and the skin by stabilizing the local pressureon the skin in the vicinity of the depth scans. For a flat opticalwindow, the pressure on the skin is distributed widely across the entireskin interface of the optical window, which is a relatively wide area.This wide distribution of pressure reduces the optical couplingefficiency of the sensor. Accordingly, the dome shape of the Petzvalsurface concentrates the pressure on the skin tissue towards the centerof the dome where the scan is taking place, which optimizes the opticalcoupling efficiency of the sensor.

According to another aspect of the present embodiment, a pedestal shapemay be incorporated onto the skin interface side of the optical window,to stabilize the local pressure on the skin in the vicinity of the depthscans by distributing the pressure along the plateau edge of thepedestal, thereby improving the optical contact.

The Petzval surface facilitates maintaining the focal point on thesurface skin and reducing the blurring of the optical axis andmaximizing the uniformity of light captured entering and exiting theskin at all points in the area scan. Using the Petzval surface, wheneverthe focused beam hits the surface of the skin, it is focused andmaximized, providing the highest efficiency of the light as well asmaintaining the same distance in depth that would be available along theoptical axis due to the skin wrapping around the Petzval surface. Thesize of the Petzval surface is a function of the focusing lens design inthe disposable apparatus. Both depth resolution and optical collectionefficiency are optimized by maintaining the focal point on the Petzvalsurface.

According to another embodiment of the present invention, a method forimproving the optical interface between a sensor and a surface of theskin includes the step of using an index matching medium at this opticalinterface, where the medium improves and stabilizes the opticalinterface and provides an optical transition for an optimal amount ofincident light from the sensor to pass through to the skin. A widevariety of mediums that can be used, each with differing opticalproperties and viscosities, such as, for example, fluids such asglycerin, saline, and mineral oil, gels, such as medical gels or a gelmoleskin, or adhesive-type materials, so long as the refractive index ofthe medium is less than the refractive index of the disposableapparatus. Preferably, the index matching medium provides a thinconformal coating on the skin and the associated disposable interface,and smoothes the optical roughness of the skin, reducing the loss ofincident light entering the skin. By using an index matching medium, apatient need not wait the 60-90 minutes for the interface of thedisposable and the skin to stabilize, but may use the OCT sensor at anygiven time by simply connecting it to the disposable optical lensapparatus adhered to the skin.

Additionally, the index matching medium smoothes out the relativelyrough surface of the skin, which may cause a scattering of the focusedbeam at the skin surface. Accordingly, the index matching medium coatsthe skin and reduces the optical roughness of the skin surface, therebyoptimizing the intensity of the light that goes into and comes out ofthe skin.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention will be more readily understood from the detaileddescription of the preferred embodiment(s) presented below considered inconjunction with the attached drawings, of which:

FIG. 1 illustrates a rotating dual wedge prism optical scanningapparatus, according to an embodiment of the present invention;

FIG. 2 illustrates a mirror based optical scanning apparatus, accordingto an embodiment of the present invention;

FIG. 3 graphically shows how the relative position of an object beingscanned by a rotating wedge prism optical scanning apparatus changes dueto the orientation of the wedge prism;

FIGS. 4A-4C illustrate the relationship between the angular velocity ofone or more wedge prisms and the depth scan rate of a sensor in relationto the scan pattern of the sensor, according to an embodiment of thepresent invention;

FIG. 5A presents a magnified view of the optical interface between anoptical window and a surface of skin;

FIG. 5B illustrates the effect of sweat and bodily fluids on the dataproduced by an optical signal;

FIG. 5C presents a magnified view of the effect of sweat and bodilyfluids on an optical interface between an optical window and a surfaceof skin;

FIG. 6A presents an optical sensor system, according to an embodiment ofthe present invention;

FIG. 6B presents an optical scanning system, according to an embodimentof the present invention;

FIG. 7A presents a Petzval surface design for a disposable optical lensapparatus, according to an embodiment of the present invention;

FIG. 7B presents a pedestal surface design for a disposable optical lensapparatus, according to an embodiment of the present invention;

FIG. 8 presents a method of using an optical scanning apparatus tomeasure blood glucose, according to an embodiment of the presentinvention;

FIG. 9 presents a method for stabilizing a scan pattern of an opticalscanning apparatus, according to an embodiment of the present invention;

FIG. 10A is a graphical illustration of varying the angular velocitiesof dual wedge prisms in an optical scanning apparatus over time; and

FIG. 10B illustrates the effect of varying the angular velocities ofdual wedge prisms in an optical scanning apparatus in comparison to thedepth scan rate of the sensor apparatus, according to an embodiment ofthe present invention.

DETAILED DESCRIPTION OF THE INVENTION

FIG. 6A presents an optical scanning apparatus system or sensor systemfor taking blood glucose measurements, according to an embodiment of thepresent invention. Specifically, the sensor system in FIG. 6A includes adual wedge prism sensor housing 616 attached to a disposable opticallens apparatus 608 with a Petzval surface 609. In FIG. 6A, sensor system601 comprises a sensor housing 614 that includes a collimator 602connected to a light source at a connecter 607, wherein the light sourceproduces a collimated light 603. An example of a connecter is afiber-optic cable. The collimated light 603 hits a fixed mirror 604,which bends the collimated light 603 to a ninety degree angle. Thecollimated light 603 passes through rotating dual wedge prisms 605 thatdeviate the angle of collimated light 603 off the optical axis of thesensor 601. The amount of deviation is based on the thickness of eachwedge prism 605 that the collimated light 603 passes through as thewedge prisms 605 rotate. The collimated light 603 then passes through afocusing lens 606, which combines the collimated light 603 intoconverged light 612, and facilitates focusing the converged light 612 tothe focal plane and focal point 611. The converged light 612 then passesthrough a disposable optical apparatus 608. The disposable apparatus 608provides an interface between the sensor and the surface of the skin 610and facilitates setting a distance from focusing lens 606 to the focalplane that is fixed at the skin surface 610 by positioning the interfaceof the skin surface 610 with the optical window 608 to the focal plane.Because the focal point 611 traces out a curved path as it deviates fromthe optical axis, attached to the bottom surface of the disposableapparatus 608 is a Petzval dome 609 that acts as an optical window andfocuses the focal point 611 onto the surface of the skin 610. As shownin FIG. 6A, the Petzval surface 609 is a separate component physicallyattached to the bottom surface of the disposable apparatus 608.Alternately, the Petzval surface 609 may be integrally formed from thesame material as the disposable apparatus 608. A data collecting device,such as a computer may connect to the sensor housing 616 via theconnector 602.

In FIG. 6B, an interferometer, an optical receiver, a demodulator, andan optical source may be miniaturized and coupled directly to the sensorhousing via the connector 607, as shown at 615, making the sensor a“sample arm” of the interferometer. Additionally, the interferometer 615may be connected to a computer 616 that downloads the sensor data andmanipulates the data to produce a blood level glucose or otherphysiological reading.

In FIG. 6A, the disposable optical lens apparatus 608, including thefocusing lens 606 and the Petzval surface 609, may be attached and lefton the skin 610 using a topical adhesive, such as, for example,cyanoacrylate or medical adhesive, such as 3M Medical Adhesive. Thesensor housing 614 then attaches to the disposable apparatus 608 atconnectors 613. When a patient has completed taking a glucose reading,the patient may remove the sensor housing 614 and leave the disposableapparatus 608 attached to the skin. Thus, for the next glucose reading,which may be at some later point in time, perhaps after a meal, thepatient need not worry about trying to place the sensor system 601 inthe same location as the previous reading in order to produce comparableresults. Instead, the patient may merely attach the sensor housing 614to the disposable apparatus 608 using connectors 613 whenever a glucosereading is desired. The disposable apparatus 608 then may be removed anddiscarded at the end of a day, for example, and replaced with a newdisposable apparatus 608 the following day. Alternately, the patient mayleave the sensor housing 614 attached to the disposable apparatus 608for an extended period of time to permit continuous blood glucosereadings.

FIGS. 7A and 7B present disposable optical lens apparatuses, accordingto an embodiment of the present invention. As shown in FIG. 7A,collimated light 603 pass through the focusing lens 606 and combine tobecome converged light 612 to pass through the disposable opticalapparatus 608. The converged light 612 focus into focal point 611 on thefocal plane. The focal plane is captured by the dome-shaped Petzvalsurface 609 attached to the bottom surface of the disposable apparatus608. The Petzval surface 609 ensures that the focal point 611 remains atthe skin interface to optimize the amount of light entering and exitingthe skin 610. FIG. 7B presents a similar design of a disposable opticalapparatus 608, but with a pedestal-shaped optical window 609, accordingto an embodiment of the present invention.

FIG. 8 presents an exemplary method of using the optical sensor system601 for blood glucose measurements. The steps of the method need not bein the sequence illustrated, and some steps may occur essentiallysimultaneously. At step S801, a patient may place or rub an indexmatching medium, such as glycerine, onto an area of skin 610 where ablood glucose reading is to be taken. Use of an index matching mediumfacilitates matching the indices of refraction between the material ofthe Petzval surface 609 with the patient's skin 610 in order to optimizethe amount of light that enters and exits the skin 610, and expeditesthe time required for the Petzval surface 609 to reach equilibrium withthe skin surface 610. For example, if the material used in the Petzvalsurface 609 has an index of refraction of 1.5 and the patient's skin 610has an index of refraction of 1.3, then without an index matching mediumsome of the focused converged light 612 entering the skin is lost due tothe lower index of refraction of the skin 610. Accordingly, not all ofthe light exits the skin 610 due to the lower index of refraction, whichcauses a loss of data. By using an index matching medium with, in thisexample, a refractive index of 1.4, the medium provides an opticaltransition for the converged light 612 between the Petzval surface 609and the skin 610, which increases the amount of light that enters andexits the skin 610. Without the index matching medium, a patient wouldhave to wait upwards of 60 to 90 minutes for the skin to produce sweatand other skin oils at the area where the disposable is placed, in orderto optimize the data collection of the sensor.

With the medium in place, at step S802, the patient may adhere thedisposable lens apparatus 608 to the area where the index matchingmedium was placed. Common adhesives such as cyanoacrylate or medicaladhesive may be used to secure the disposable apparatus 608 to the skin610. Once the patient feels that the disposable apparatus 608 is secure,at step S803, the patient couples the sensor housing 614 to thedisposable apparatus 608 using the connectors 613.

At step S804, sensor diagnostics verify that a threshold trigger of 45dB has been pre-set to normalize the scans and resolve for variations inthe optical path lengths of the scans produced by the rotating wedgeprisms 605 and, accordingly, the change in the thickness of each wedgeprism 605 during the rotations. At step S805, sensor diagnostics verifythat the angular velocity of each wedge prism 605 has been pre-set to avalue such that the lateral position of each focused scan spot movesless than 10× the diameter of the focused scan spot during the dataacquisition of the depth scan. For example, if focused scan spot sizehas a diameter of 20 microns, then the angular velocity is set to avalue such that the focused beam 611 does not move laterally more than200 microns during the depth scan. By setting the angular velocity ofeach wedge prism 605 to such a value, the distortion in the depth scaleof each scan produced by the change in thickness of the wedge prism 605as it rotates is minimized. The threshold trigger, depth scan rate andangular velocities are presets that may be optimized and built into thesensor system 601.

At step S806, the patient sets the sensor system 601 to begin scanningthe skin 610. Since a threshold trigger was set at 45 dB in step S804,the sensor system 601 will not accumulate scan data until the intensityof the optical signal produced by the sensor system 601 reaches a valueof 45 dB. Preferably, the threshold is above the highest noise peakproduced by the signal but at least 10 dB lower than the intensity peakat the interface between the skin 610 and the disposable apparatus 614.

Once the sensor system 601 has completed taking multiple scans,preferably around 1500 scans, at step S807, the sensor housing 614 maybe removed from the disposable apparatus 608, or, alternately, thesensor housing 614 may remain and begin to take another glucose reading.The disposable apparatus 608 remains adhered to the patient's skin 610.The scan data then is manipulated by computer 616 connected to theinterferometer 615. Because the threshold trigger was used, all thescans taken begin at a signal intensity of 45 dB, which is equivalent toTime 0, and accordingly, at step S808, the scans are averaged to reducethe speckle associated with the sensor 601. At step S809, the averagedscan data is manipulated using algorithms, such as those described inU.S. Provisional Applications Nos. 60/671,007 and 60/671,285, to deriveblood glucose levels. At any later time, such as after a meal, thepatient may reattach the sensor housing 614 to the disposable apparatus608 to take another glucose measurement.

Alternately, the sensor system 601 may be designed to not use athreshold trigger setting at S804, and may normalize the scans once thedata has been acquired. For example, once the sensor completes a glucosereading at step S807, computer 616 of the sensor system 601 may apply apeak locating algorithm such as, for example, Gaussian peak fitting, tothe first scan to locate the first peak, at step S810. Once step S810has been completed, the peak locating algorithm is applied to eachsuccessive scan, as shown at step S811. At step S812, the successivescans are normalized in depth against the first scan by essentiallydesignating the location of each peak as at Time 0, in order to averagethe scans together. Thus, any distortion in the optical path length dueto the change in the thickness of the wedge prisms 605 as they rotate isremoved.

FIG. 9 presents an exemplary method for stabilizing the scan pattern ofsensor 601 and is discussed in conjunction with FIG. 10A, which is agraphical illustration of varying the angular velocities of the dualwedge prisms 605 of sensor system 601. When using the sensor system 601to take a blood glucose measurement, the first wedge prism 605 beginsrotating at a rate of 2.1 revolutions per second (“rps”), which isequivalent to 2.1 Hz, at step S901, as shown at 1001 in FIG. 10A.Similarly, at step S902, the second wedge prism 605 begins rotating at arate of 1.3 Hz, as shown at 1002, where 2.1 Hz and 1.3 Hz are notintegrals of each other. The sensor system 601 then begins to performdepth scans at a rate of 30 Hz, at step S903. An integral of 30 Hz is 2Hz (i.e., 2 multiplied by 15 equals 30). Additionally, another integralof 30 Hz is 1.5 Hz (i.e., 1.5 multiplied by 20 equals 30 Hz). Thus,although the wedge prisms 605 begin to rotate at rates that arenon-integrals of 30 Hz, if the angular velocities 1001, 1002 of bothwedge prisms 605 remain at 2.1 Hz and 1.3 Hz, the angular velocities maydrift towards 2.0 Hz and 1.5 Hz, thereby becoming integrals of 30 Hz,and preventing conformal coverage of the scan pattern area of the skin610.

To prevent the angular velocities from becoming integrals of the depthscan rate and remaining at the integral rates, both angular velocities1001 and 1002 of the wedge prisms 605 are varied over time, in relationto the depth scan rate and in relation to each wedge prism 605, as shownin FIG. 10A. At step S904, the angular velocity 1001 of the first wedgeprism 605 is varied as the sensor system 601 continues to perform depthscans. In FIG. 10A, the angular velocity 1001 of the first wedge prism605 is sinusoidal, oscillating from 2.5 Hz to 1.7 Hz, over a period of6500 milliseconds, or 6.5 seconds. At step S905, the angular velocity1002 of the second wedge prism 605 is varied independent of the angularvelocity 1001 of the first wedge prism 605, as shown in FIG. 10A. InFIG. 10A, the angular velocity 1002 of the second wedge prism 605 issinusoidal, oscillating from 1.55 Hz to 1.1 Hz, over a period of 5250milliseconds, or 5.25 seconds. Thus, although the angular velocities ofboth wedge prisms 605 may hit a harmonic of 30 Hz during the variation,the angular velocities only remain an integral of 30 rpm for one or twodepth scans before the velocities change, thereby minimizing the loss ofdepth scan data due to the angular velocities being integrals of thedepth scan rate. The result is a random, conformal mapping of thescanned surface area of the skin 610 with minimal overlapping within theresults, as shown at step S906.

FIG. 10B illustrates the results of varying the angular velocities ofthe wedge prisms 605 over time with respect the depth scan rate ofsensor system 601 and with respect to each wedge prism 605. Byminimizing the potential for a harmonic phase to be created between thedepth scan rate and the angular velocities of the wedge prisms 605,conformal coverage of the area of skin 610 scanned is optimized, witheach dot representing a position of an individual depth scan on the skin610.

While the present invention has been described with respect to what ispresently considered to be the preferred embodiments, it is to beunderstood that the invention is not limited to the disclosedembodiments. To the contrary, the invention is intended to cover variousmodifications and equivalent arrangements included within the spirit andscope of the appended claims. The scope of the following claims is to beaccorded the broadest interpretation so as to encompass all suchmodifications and equivalent structures and functions.

1. A system for creating a stable and reproducible optical interface inan optical sensor system for measuring blood glucose levels inbiological tissue comprises: an optical interferometer sensor utilizinga beam of light; an interferometer connected to the optical sensor; acomputer connected to an optical receiver associated with theinterferometer; and a disposable optical lens apparatus connected to theoptical sensor, wherein the disposable apparatus comprises: a focusinglens; and an optical window connected to the focusing lens; wherein thedisposable apparatus is attachable to a surface area of biologicaltissue by an adhesive.
 2. The system of claim 1, wherein the opticalsensor utilizes at least one rotating wedge prism to deviate an angle ofthe beam of light from an optical axis of the system.
 3. The system ofclaim 1, wherein the optical sensor utilizes at least one oscillatingangled mirror to deviate an angle of the beam of light from an opticalaxis of the system.
 4. The system of claim 1, wherein the optical windowincludes a Petzval surface to interface with the surface area of thebiological tissue.
 5. The system of claim 4, wherein the Petzval surfaceis in the shape of a dome.
 6. The system of claim 1, wherein the opticalwindow is in the shape of a pedestal.
 7. A method of using an opticalsystem with an optical interferometer sensor and a disposable opticalapparatus for measuring blood glucose levels in biological tissuecomprises the steps of: attaching the disposable apparatus to a surfacearea of biological tissue using an adhesive; coupling the optical sensorto the disposable apparatus; and taking a blood glucose reading usingthe optical sensor.
 8. The method of claim 7, further comprising thesteps of: removing the optical sensor from the disposable apparatus; andleaving the disposable apparatus on the surface area of biologicaltissue for later use.
 9. The method of claim 7, further comprising thestep of placing an index matching medium between the disposableapparatus and the surface area of the biological tissue.
 10. A methodfor resolving variations in an optical path length of an optical sensorsystem having at least one wedge prism for measuring blood glucoselevels in biological tissue comprises the following steps: taking aplurality of scans using the optical sensor system; locating a firstpeak in a first scan of the plurality of scans; locating a first peak ineach subsequent scan of the plurality of scans; normalizing each firstpeak in each subsequent scan against the first peak in the first scan;and averaging the normalized scans to produce a resulting averaged scan.11. The method of claim 10, further comprising the step of using a peaklocating algorithm to locate each first peak.
 12. A method for resolvingvariations in an optical path length of an optical sensor system havingat least one wedge prism for measuring blood glucose levels inbiological tissue comprises the following steps: determining a thresholdtrigger; and setting the optical sensor system to begin acquiring dataonce the threshold trigger is reached.
 13. The method of claim 12,further comprising the step of setting the threshold trigger to a signalintensity value of at least 10 decibels below a first peak intensityvalue associated with a signal produced by the optical sensor system,wherein the first peak intensity value corresponds to an opticalinterface between the optical sensor system and the biological tissue.14. The method of claim 13, further comprising the step of setting thepeak threshold trigger to a signal intensity value above a highest noisepeak associated with the signal.
 15. The method of claim 12, wherein thestep of determining further comprises setting the threshold trigger to asignal intensity value that relates to a specific structural feature.16. The method of claim 15, wherein the specific feature corresponds toan optical interface between the optical sensor system and thebiological tissue.
 17. The method of claim 12, further comprising thestep of setting an angular velocity of at least one wedge prism to avalue such that a lateral position of a beam of light utilized by theoptical sensor system on the biological tissue moves a distance that isless than a value equal to a diameter of the beam of light multiplied bya factor of ten, during a single depth scan.
 18. A method for minimizingdistortion of a depth scale of an optical sensor system having at leastone wedge prism for measuring blood glucose levels in biological tissuecomprises the step of setting an angular velocity of at least one wedgeprism to a value such that a lateral position of a beam of lightutilized by the optical sensor system on the biological tissue moves adistance that is less than a value equal to a diameter of the beam oflight multiplied by a factor of ten, during a single depth scan.
 19. Amethod for stabilizing a scan pattern of an optical sensor system havingat least one wedge prism for measuring blood glucose levels inbiological tissue comprises the step of setting an angular velocity ofthe at least one wedge prism to a non-integral value of a depth scanrate associated with the optical sensor system.
 20. The method of claim19, further comprising the step of varying the angular velocity of theat least one wedge prism while the optical sensor system is in use. 21.The method of claim 19, further comprising the step of varying theangular velocity of the at least one wedge prism with respect to asecond angular velocity associated with a second wedge prism during atotal time of use of the optical sensor system, wherein the angularvelocity is a non-integral value of the second angular velocity.
 22. Amethod for stabilizing a scan pattern of an optical sensor system havinga plurality of wedge prisms for measuring blood glucose levels inbiological tissue comprises the steps of: setting a first angularvelocity associated with a first wedge prism to a harmonic integral of asecond angular velocity associated with a second wedge prism, whereinthe first angular velocity is at least one order of magnitude greaterthan the first angular velocity; and setting the first angular velocityand the second angular velocity to non-harmonic integrals of a depthscan rate associated with the optical sensor system.
 23. A method foroptimizing an amount of light entering and exiting an area of biologicaltissue, wherein the amount of light is associated with an optical sensorsystem placed on the area of biological tissue in order to measure bloodglucose levels, comprising the steps of: coupling a Petzval surface to abottom surface of the optical sensor system; placing the optical sensorsystem on the surface area of biological tissue, such that the tissuewraps around the Petzval surface; directing a beam of light through afocusing lens and the optical window, into the Petzval surface; andmaintaining a focal point of the beam of light at an interface of thePetzval surface and the area of biological tissue.
 24. The method ofclaim 23, further comprising the step of injecting an index matchingmedium into the interface of the Petzval surface and the area ofbiological tissue, wherein the medium automatically matches a refractiveindex of the area of biological tissue with a refractive index of thePetzval surface.